1. Field of the Invention
The present invention relates to ophthalmic instruments that are used to examine or treat the eye, including ophthalmic examination instruments (such as phoropters and autorefractors) that measure and characterize the aberrations of the human eye in order to prescribe compensation for such aberrations via lens (such as glasses or contact lens) or surgical procedure (such as laser refractive surgery), in addition to ophthalmic imaging instruments (such as fundus cameras, corneal topographers, retinal topographers, corneal imaging devices, and retinal imaging devices) that capture images of the eye.
2. Summary of the Related Art
The optical system of the human eye has provided man with the basic design specification for the camera. Light comes in through the cornea, pupil and lens at the front of the eye (as the lens of the camera lets light in). This light is then focused on the inside wall of the eye called the retina (as on the film in a camera). This image is detected by detectors that are distributed over the surface of the retina and sent to the brain by the optic nerve which connects the eye to the brain (as film captures the image focused thereon).
FIG. 1 shows a horizontal cross section of the human eye. The eye is nearly a sphere with an average diameter of approximately 20 mm. Three membranes—the cornea and sclera outer cover, the choroid and the retina—enclose the eye. The cornea 3 is a though transparent tissue that covers the anterior surface of the eye. Continuous with the cornea 3, the sclera 5 is an opaque membrane that encloses the remainder of the eye. The choroid 7 lies directly below the sclera 5 and contains a network of blood vessels that serves as the major source of nutrition to the eye. At its anterior extreme, the choroid 7 includes a ciliary body 9 and an iris diaphragm 11. The pupil of the iris diaphragm 11 contracts and expands to control the amount of light that enters the eye. Crystalline lens 13 is made up of concentric layers of fibrous cells and is suspended by fibers 15 that attach to the ciliary body 9. The crystalline lens 13 changes shape to allow the eye to focus. More specifically, when the ciliary muscle in the ciliary body 9 relaxes, the ciliary processes pull on the suspensory fibers 15, which in turn pull on the lens capsule around its equator. This causes the entire lens 13 to flatten or to become less convex, enabling the lens 13 to focus light from objects at a far away distance. Likewise, when the ciliary muscle works or contracts, tension is released on the suspensory fibers 15, and subsequently on the lens capsule, causing both lens surfaces to become more convex again and the eye to be able to refocus at a near distance. This adjustment in lens shape, to focus at various distances, is referred to as “accommodation” or the “accommodative process” and is associated with a concurrent constriction of the pupil.
The innermost membrane of the eye is the retina 17, which lies on the inside of the entire posterior portion of the eye. When the eye is properly focused, light from an object outside the eye that is incident on the cornea 3 is imaged onto the retina 17. Vision is afforded by the distribution of receptors (e.g., rods and cones) over the surface of the retina 17. The receptors (e.g., cones) located in the central portion of the retina 17, called the fovea 19 (or macula), are highly sensitive to color and enable the human brain to resolve fine details in this area. Other receptors (e.g., rods) are distributed over a much larger area and provides the human brain with a general, overall picture of the field of view. The optic disc 21 (or the optic nerve head or papilla) is the entrance of blood vessels and optic nerves from the brain to the retina 17. The inner part of the posterior portion of the eye, including the optic disc 21, fovea 19 and retina 17 and the distributing blood vessels in called the ocular fundus 23. Abnormalities in the cornea and crystalline lens and other portions of the eye contribute to refractive errors (such as defocus, astigmatism, spherical aberrations, and other high order aberrations) in the image captured by the retina.
A phoropter (or retinoscope) is an ophthalmic instrument that subjectively measures the refractive error of the eye. A typical phoropter consists of a pair of housings in which are positioned corrective optics for emulating the ophthalmic prescription required to correct the vision of the patient whose eyes are being examined. Typically, each housing contains sets of spherical and cylindrical lenses mounted in rotatable disks. The two housings are suspended from a stand or wall bracket for positioning in front of the patient's eyes. Further, in front of each refractor housing a number of accessories are mounted, typically on arms, so that they may be swung into place before the patient's eyes. Typically, these accessories include a variable power prism known as a Risley prism, Maddox rods, and a cross cylinder for performing the Jackson cross cylinder test. In determining a patient's distance prescription, the patient views a variety of alpha numeric characters of different sizes through various combinations of the spherical and/or cylindrical lenses supported in the refractor housings until the correct prescription is emulated. The characters, which are typically positioned 6 meters away, may be on a chart or may be projected on a screen by an acuity projector. For near vision testing the same procedure is repeated, expect that the alpha numeric characters viewed by the patient are positioned on a bracket 20 to 65 centimeters in front of the refractor housing. The cross cylinder is used to refine the power and axis position of the cylindrical component of the patient's prescription. The cross cylinder is a lens consisting of equal power plus and minus cylinders with their axes 90 degrees apart. It is mounted in a loupe for rotation about a flip axis which is midway between the plus and minus axes.
An autorefractor is an ophthalmic instrument that quantitatively measures the refractor errors of the eye. Light from an illumination source (typically an infra-red illumination source) is directed into the eye of the patient being examined. Reflections are collected and analyzed to quantitatively measure the refractive errors of the eye.
Conventional phoropters and autorefractors characterize the refractive errors of the eye only in terms of focal power (typically measured in diopter) required to compensate for such focal errors; thus, such instruments are incapable of measuring and characterizing the higher order aberrations of the eye, including astigmatism and spherical aberration. Examples of such devices are described in the following U.S. Pat. Nos. 4,500,180; 5,329,322; 5,455,645; 5,629,747; and 5,7664,561.
Instruments have been proposed that utilize wavefront sensors to measure and characterize the high order aberrations of the eye. For example, U.S. Pat. No. 6,007,204, to Fahrenkrug et al. discloses an apparatus for determining refractive aberrations of the eye wherein a substantially collimated beam of light is directed to the eye of interest. This collimated light is focused as a secondary source on the back of the eye, thereby producing a generated wavefront that exits the eye along a return light path. A pair of conjugate lenses direct the wavefront to a microoptics array of lenslet elements, where incremental portions of the wavefront are focuses onto an imaging substrate. Deviation of positions of the incremental portions relative to a known zero or “true” position (computed by calculating the distance between the centroids of spots formed on the imaging substrate by the lenslet array) can be used to compute refractive error relative to a known zero or ideal diopter value. Because the optical power at the lenslet does not equal the optical power of the measured eye, the optical power of the lenslet is corrected by the conjugate lens mapping function to interpolate the power of the eye. This refractive error is reported to the user of the apparatus through an attached LCD.
In U.S. Pat. Nos. 5,777,719; 5,949,521; and 6,095,651, Williams and Liang disclose a retinal imaging method and apparatus that produces a point source on a retina by a laser. The laser light reflected from the retina forms a distorted wavefront at the pupil, which is recreated in the plane of a deformable mirror and a Shack-Hartmann wavefront sensor. The Shack-Hartmann wavefront sensor includes an array of lenslets that produce a corresponding spot pattern on a CCD camera body in response to the distorted wavefronts. Phase aberrations in the distorted wavefront are determined by measuring spot motion on the CCD camera body. A computer, operably coupled to the Shack-Hartmann wavefront sensor, generates a correction signal which is fed to the deformable mirror to compensate for the measured phase aberrations. After correction has been achieved via the wavefront sensing of the reflected retinal laser-based point source, a high-resolution image of the retina can be acquired by imaging a krypton flash lamp onto the eye's pupil and directing the reflected image of the retina to the deformable mirror, which directs the reflected image onto a second CCD camera body for capture. Examples of prior art Shack-Hartmann wavefront sensors are described in U.S. Pat. Nos. 4,399,356; 4,725,138, 4,737,621, and 5,529,765; each herein incorporated by reference in its entirety.
Notably, the apparatus of Fahrenkrug et al. does not provide for compensation of the aberrations of the eye. Moreover, the apparatus of Fahrenkrug et al. and the apparatus of Williams and Liang do not provide a view of the compensation of the aberrations to the eye. Thus, the patient cannot provide immediate feedback as to the accuracy of the measurement; and must wait until compensating optics (such as a contact lens or glasses that compensate for the measured aberrations) are provided in order to provide feedback as to the accuracy of the measurement. This may lead to repeat visits, thereby adding significant costs and inefficiencies to the diagnosis and treatment of the patient.
In addition, the wavefront sensing apparatus (i.e., the lenslet array and imaging sensor) of Fahrenkrug et al. and of Williams and Liang are susceptible to a dot crossover problem. More specifically, in a highly aberrated eye, the location of spots produced on the imaging sensor may overlap (or cross). Such overlap (or crossover) introduces an ambiguity in the measurement that must be resolved, or an error will be introduced.
In addition, the signal-to-noise ratio provided by traditional Hartmann sensing techniques in measuring the aberrations of the human eye is limited, which restricts the potential usefulness of ophthalmic instruments that embody such techniques in many real-world ophthalmic applications. More specifically, the basic measurement performed by any Hartmann wavefront sensor is the determination of the locations of the Hartmann spots. Traditionally, this has been done by calculating the centroid of the illumination in a pixel subaperture defined around each spot.
Centroid calculation is conceptually very simple. To calculate the centroid of the light distribution in the x-direction, weights are assigned to each column of pixels in the pixel subaperture and the measured intensity for each pixel in the pixel subaperture is multiplied by the weight corresponding to the column of the given pixel and summed together. If the weights vary linearly with the distance of the column from the center of the pixel subaperture, this sum will be a measure of the x-position of the light distribution. The sum needs to be normalized by dividing by the sum of the unweighted intensities. To calculate the centroid of the light distribution in the y-direction, weights are assigned to each row of pixels in the pixel subaperture and the measured intensity for each pixel in the pixel subaperture is multiplied by the weight corresponding to the row of the given pixel and summed together. If the weights vary linearly with the distance of the column from the center of the pixel subaperture, this sum will be a measure of the y-position of the light distribution. The sum needs to be normalized by dividing by the sum of the unweighted intensities. Such centroid calculation may be represented mathematically as follows:             x      c        =                            ∑          i                ⁢                              ∑            j                    ⁢                                    w              j                        *                          I              ij                                                            ∑          i                ⁢                              ∑            j                    ⁢                      I            ij                                          y      c        =                            ∑          i                ⁢                              ∑            j                    ⁢                                    w              i                        *                          I              ij                                                            ∑          i                ⁢                              ∑            j                    ⁢                      I            ij                              where i and j identify the rows and columns, respectively, of the pixel subaperture; wi and wj are the weights assigned to given rows and columns, respectively, of the pixel subaperture; and Iij is the intensity of a given pixel in row i and column j of the pixel subaperture.
This “center-of-light” measurement is analogous to the usual center-of-mass calculation. FIG. 2 shows a one dimensional representation of the intensity distribution on a row of detector pixels and a set of weights. These weights are simply the distance of the center of the pixel from the center of the pixel subaperture in units of pixel spacing.
However, centroid calculation is disadvantageous because it is susceptible to background noise and thus may be unacceptable in many real-world environments where background noise is present. FIG. 2 reveals these shortcomings. Note that the highest weights are applied to pixels farthest from the center. Note also that, typically, there is very little light in these regions. This means that the only contribution to these highly weighted pixels comes from background light and noise. Because of the high weight, these pixels adversely affect the accuracy of the measurement. As the size of the pixel region of that measures such spot motion is increased to provide greater tilt dynamic range, the noise problem is made worse by increasing the number of pixels that usually have no useful signal.
An even larger problem stems from the centroid algorithms sensitivity to residual background signal. Consider a pixel region that is 10×10 pixels in size. Typically, a given spot will occupy less than 10 of those pixels. Suppose there is a residual background signal that produces, per pixel, 1% of the signal from the given spot. Because it is present in all 100 pixels, its contribution to the total signal is equal to that of the spot. Even if this background is perfectly uniform, when the centroid is normalized by the total signal, that divisor will be twice its true size. This will make the calculated centroid half is correct value. If the background is not uniform, its effect on the centroid can easily overwhelm that of the spot. Thus, the susceptibility of the centroid algorithm to background noise makes it unacceptable in many real-world environments where such background noise is present.
Thus, there is a great need in the art for improved ophthalmic instruments that measure and characterize the aberrations of the human eye in a manner that avoids the shortcomings and drawbacks of prior art ophthalmic instruments.